The field of the invention is medical imaging and more particularly medical imaging with a combined CT-PET system.
Throughout this specification, in the interest of simplifying this explanation, a clinical region of a patient to be imaged will be referred to generally as a xe2x80x9cregion of interestxe2x80x9d and prior art and the invention will be described with respect to a hypothetical region of interest. In addition, the phrase xe2x80x9ctranslation axisxe2x80x9d will be used to refer to an axis along which a patient is translated through an imaging system during data acquisition.
The medical imaging industry has developed many different types of imaging systems that are useful for diagnostic purposes. Two of the more widely used systems include computerized tomography (CT) systems and positron emission tomography (PET) systems.
In CT systems, an X-ray-source projects a fan-shaped X-ray beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system, termed the xe2x80x9cCT imaging plane.xe2x80x9d The X-ray beam passes through a region of interest, such as the torso of a patient, and impinges upon an array of radiation detectors. The intensity of the transmitted radiation is dependent upon the attenuation of the X-ray beam by the region of interest and each detector produces a separate electrical signal that is a measurement of the beam attenuation. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile. A group of X-ray attenuation measurements from the detector array at a given angle is referred to as a xe2x80x9cviewxe2x80x9d and a xe2x80x9cscanxe2x80x9d of the object comprises a set of views made at different angular orientations during one revolution of the X-ray source. Using various data collection and manipulation techniques CT data can be used to generate two and three dimensional images of the region of interest.
Unlike CT systems that rely on an external X-ray source to generate image data, PET systems rely on an energy source that resides within a region of interest. To this end, positrons are positively charged electrons which are emitted by radionuclides that have been prepared using a cyclotron or other device. The radionuclides most often employed in diagnostic imaging are fluorine-18, carbon-11, nitrogen-13 and oxygen-15. Radionuclides are employed as radioactive tracers called xe2x80x9cradiopharmaceuticalsxe2x80x9d by incorporating them into substances such as glucose or carbon dioxide.
To use a radiopharmaceutical in PET imaging, the radiopharmaceutical is administered, typically by injection, to a patient and accumulates in an organ, vessel or the like, which is to be imaged. It is known that specific radiopharmaceuticals become concentrated within certain organs and tumors or, in the case of a vessel, that specific radiopharmaceuticals will not be absorbed by a vessel wall. Thus, to image a specific region of interest, a radiopharmaceutical known to accumulate either within the region of interest, within an organ that resides in the region of interest or within a fluid that passes through the region of interest can be selected. The process of concentrating often involves processes such as glucose metabolism, fatty acid metabolism and protein synthesis.
After the radiopharmaceutical becomes concentrated within a region of interest and while the radionuclides decay, the radionuclides emit positrons. The positrons travel a very short distance before they encounter an electron and, when the positron encounters an electron, the positron is annihilated and converted into two photons, or gamma rays. This annihilation event is characterized by two features which are pertinent to medical imaging and particularly to medical imaging using positron emission tomography (PET). First, each gamma ray has an energy of essentially 511 keV upon annihilation. Second, the two gamma rays are directed in substantially opposite directions.
In PET imaging, if the general locations of annihilations can be identified in three dimensions, a three dimensional image of a region of interest can be reconstructed for observation. To detect annihilation locations, a PET camera is employed. An exemplary PET camera includes a plurality of detectors and a processor which, among other things, includes coincidence detection circuitry. Each time an approximately 511 keV positron impacts a detector, the detector generates an electronic signal or pulse which is provided to the processor coincidence circuitry.
The coincidence circuitry identifies essentially simultaneous pulse pairs which correspond to detectors which are essentially on opposite sides of the imaging area. Thus, a simultaneous pulse pair indicates that an annihilation has occurred on a straight line between an associated pair of detectors. Over an acquisition period of a few minutes millions of coincidence events are recorded, each coincidence event is associated with a unique detector pair. After an acquisition period during which coincidence data is collected from every angle about an imaging area, recorded coincidence data can be used via any of several different well known procedures to construct images of radionuclide concentration in the region of interest. In the case of PET systems, PET data can be collected simultaneously from a volume within an object of interest so that a 3D image can be generated.
As is the case in virtually all imaging systems, one measure of the value of a PET system is throughput. To this end, in a radiology department the number of images generated is generally related to profitability with greater numbers of images translating into greater profitability. Thus, PET acquisition systems are generally designed to collect required imaging data rapidly. For this reason, one well accepted PET configuration is generally referred to as a full ring system which, as its label implies, includes a plurality of detector segments arranged to form an annular detector surface about an imaging area such that the system detects annihilation photons from many angles at a time. Hereinafter, for the purposes of this explanation a full ring detector system will be assumed unless indicated otherwise.
Each of the different imaging modalities typically has uses for which it is particularly advantageous. For example, CT systems that employ X-rays are useful for generating anatomical images of bone and the like while PET systems are useful for generating functional images corresponding to dynamic occurrences such as metabolism and the like.
For various reasons, in some diagnostic applications, it is advantageous to collect both CT and PET data corresponding to the same clinical region of interest. For instance, CT data may be used to compensate for inaccuracies in PET imaging data. To this end, one of the problems with PET imaging techniques is that gamma ray absorption and scatter by portions of the region of interest between the emitting radiopharmaceutical and a detector distort collected data and hence resultant images. One solution for compensating for gamma ray attenuation is to assume uniform positron attenuation throughout the region of interest. That is, the region of interest is assumed to be completely homogenous in terms of radiation attenuation with no distinction made for bone, soft tissue, lung, etc. This assumption enables attenuation estimates to be made based on the surface contour of the region of interest. Unfortunately, typical regions of interest do not cause uniform radiation attenuation and therefore a uniform attenuation assumption is generally inaccurate.
According to several methods, instead of assuming uniform attenuation characteristics throughout the region of interest, CT transmission data is collected for the entire region of interest and is used to accurately determine attenuation characteristics at every point throughout the region of interest. Thereafter, the PET emission data is corrected as a function of the CT attenuation map to generate more accurate PET images.
As another instance where it is advantageous to generate both CT and PET data for a region of interest, sometimes it is advantageous to generate images that include both anatomical and functional characteristics. To this end, one solution has been to sequentially use separate imaging systems to gather both functional and static imaging data sets and then combine those sets or corresponding images to generate unified functional/static images. For example, a CT system may be used to generate an anatomical CT image and subsequently a PET system may be used to generate a functional PET image, the two images being combined thereafter to generate the unified image.
Unfortunately, where both CT and PET imaging data have to be acquired for a single region of interest, several configuration and processing problems have to be overcome. First, after functional and anatomical image data has been collected, there has to be some way to align the functional and anatomical images so that the unified image precisely reflects relative anatomic juxtapositions. To this end, in some cases, fiducial markers have been employed. For example, a metallic button with a positron emitter can be placed on the surface of a patient""s skin Which is detectable by both the CT and PET systems. By aligning the marker in the resulting images the images can be aligned.
Second, where two separate imaging configurations are employed, a patient has to be moved from one configuration to the next between acquisition sessions. Movement increases the likelihood that the patient""s positions during the two imaging sessions will change thus tending to reduce the possibility of accurate alignment (i.e., relative positions of organs or the like could change during movement). The possibility of misalignment is exacerbated by the fact that often imaging session schedules will not allow both CT and PET imaging processes to be performed during the same day. Thus, overall diagnostic value of the resulting unified image can be reduced appreciably through movement between acquisition periods.
Third, where two separate imaging systems are employed to obtain imaging data during two separate acquisition periods, the time required to acquire needed data is relatively long and hence throughput (i.e., the number of imaging sessions that can be performed over the course of a given period) is reduced.
One solution to eliminate the need to move patients between acquisition systems is to provide a dual CT-PET imaging system. To this end, one general type of dual imaging system includes, in effect, a CT system mounting in some fashion to a PET system so that data is sequentially collected for the region of interest, first the PET or CT data and then, after table and region of interest translation along a translation axis, the CT or PET data. These types of systems are better than two separate acquisition systems because they generally minimize patient movement and therefore facilitate data alignment.
While better than separate acquisition systems, these dual systems have several shortcomings such as bore or imaging area length as each of the adjacent system requires its own imaging area. In addition to requiring additional space, long bore length often increases patient discomfort as such systems are generally psychologically intimidating (i.e., many patients become nervous when placed in a long bore) and often result in additional patient movement. Moreover, these dual systems require two separate detector assemblies, one for CT and one for PET data acquisition. Furthermore, these types of systems still require two separate and sequential imaging periods to collect data.
One other solution to eliminate patient movement while still enabling acquisition of both CT and PET data has been to provide a radiation transmission source within the imaging area between facing PET detectors that transmits radiation toward one of the PET detectors. One exemplary system (hereinafter xe2x80x9cthe Saoudi systemxe2x80x9d) of this type has been described in an article entitled xe2x80x9cA Novel APD-Based Detector Module for Multi-Modality PET/SPECT/CT Scannersxe2x80x9d by A. Saoudi and R. Lecomte that was published in the IEEE 1999 publication. The Saoudi system including a full ring detector constructed to receive and differentiate both emission and transmission data and a transmission X-ray source. The X-ray source is positioned inside the ring detector, presumably on a track of some type, for rotation about the internal surface of the detector to direct a fan beam across an imaging area within the detector bore.
To operate the Saoudi system, after a radionuclide has accumulated within an organ that resides in a region of interest, a patient is positioned on a support table with the region of interest inside the imaging area. Thereafter the transmission source is turned on to direct the fan beam across the imaging area through the region of interest and toward a portion of the detector on the opposite side of the imaging area. The section of the detector subtended by X-rays from the transmission source acquires both annihilation photons from the radionuclide and also transmitted X-rays from the source while other sections of the detector that are not subtended by the X-rays collect only annihilation photon data. The source is rotated about the imaging area during a data acquisition process so that transmission data is collected from every angle about the imaging area. Sorting circuitry differentiates between the different energies of the annihilation photons and the X-rays and thus two separate sets of data, an emission set and a transmission set, are acquired.
The Saoudi system is better than the previous solutions but also has several disadvantages. First, when the source is between a detector segment and the concentrated radionuclide (i.e., the organ being imaged), the source blocks emission data from reaching the detector segment and some of the PET data is lost which either results in a less accurate image or increases the time necessary to collect required PET data.
Second, because Saoudi teaches placement of the fan beam X-ray source inside the detector and essentially inside the imaging area, the size of the imaging area has to be increased to accommodate a patient and still provide sufficient space for the source to rotate therein between the patient and the detector. A larger imaging are also requires a larger full ring detector which increases costs appreciably.
U.S. Pat. No. 5,600,145 (hereinafter xe2x80x9cthe ""145 patentxe2x80x9d) which issued on Feb. 4, 1997 and is entitled xe2x80x9cEmission/Transmission Device for use with a Dual Head Nuclear Medicine Gamma Camera with the Transmission Source Located Behind the Emission Collimatorxe2x80x9d describes one attempt to reduce the size of a combined CT-PET system. To this end, the ""145 patent system teaches opposing first and second emission/transmission detectors and a line transmission source having a length that is substantially similar to the length of the first detector. The line source is mounted between the first detector and a collimator corresponding to the first detector so that the source generates a line of radiation directed across an imaging area between the first and second detectors and toward the second detector with the source length traversing across the first detector length and for movement across a first detector width perpendicular to the detector length.
In operation, after a radiopharmaceutical has accumulated in an organ that is within a region of interest, the region of interest is positioned between the detectors, the line source is turned on and is moved across the width of the first detector and the second detector collects both emission and transmission data corresponding to the radiopharmaceutical and the source, respectively while the portion of the first detector that is not blocked by the source collects emission data. The detectors (including the X-ray source) are rotated about the region of interest to collect data from all angles and, after acquisition, a sorter sorts the data by energy levels into emission and transmission data.
Here, the ""145 patent teaches that by placing the line source between the detector surface and the detector collimator the distance between the emitting positron source and the detector surface can be minimized as there is no need for an additional collimator for the line source. Nevertheless, the line source, like the source in the Saoudi reference, increases the positron source-detector distance. In addition, it is unclear how the teachings of the ""145 patent could be applied in the case of a full ring detector to expedite emission data acquisition.
It has been recognized that, where a PET detector assembly includes first and second detector segments on opposite sides of an imaging area, an X-ray source can be placed outside the imaging area and adjacent one of the detector segments where the source generates an X-ray directed across the imaging area and toward the other segment of the detector such that the detector assembly and the source can be used to essentially simultaneously collect both transmission and emission data thereby overcoming many of the problems discussed above.
Consistent with the above description, the invention includes a photon imaging apparatus comprising oppositely facing first and second emission/transmission detector segments disposed on opposite sides of an imaging area and defining the imaging area there between and a radiation source disposed adjacent and outside the imaging area and adjacent at least one of the detector segments, the source generating a fan beam of radiation that emanates from a focal point and juxtaposed such that the fan beam is directed along a trajectory through the imaging area and toward the other of the detector segments, wherein the segments collect both emission and transmission radiation.
The invention also includes a method for collecting both emission and transmission imaging data, sequentially or concurrently, corresponding to a region of interest within a patient where a radiopharmaceutical has accumulated within the region of interest and is generating emission radiation. The method comprises the steps of placing the region of interest within an imaging area between first and second photon detector segments, directing a fan beam of transmission radiation from a focal spot adjacent and outside the imaging area along a trajectory through the imaging area and toward at least one of the detector segments and collecting both the emission and transmission radiation via the first and second detector segments.
Moreover, the invention also includes a combined emission-transmission imaging apparatus, the apparatus comprising a full ring photon detector assembly including detector segments that are arranged to form an annular detector surface about an imaging area and radiation source disposed adjacent and outside the imaging area and adjacent a first side of the detector assembly, the source generating a fan beam of radiation that emanates from a focal point and juxtaposed such that the fan beam is directed along a trajectory through the imaging area and toward the detector surface on an opposite side of the imaging area, the source controllable to alter the position of the focal point and hence the trajectory of the fan beam such that the fan beam is directable across the imaging area along various trajectories and toward various segments of the detector surface, wherein the segments collect both emission and transmission radiation.
To reduce patient exposure to X-rays, in at least one embodiment of the invention, a CT source collimator is provided that collimates the CT X-ray beam in a desired fashion. To this end, the collimator may include two jaws having either straight or curved collimating edges separated along the Z-axis. The collimator may be stationary or, in the alternative, may be rotated along with the point at which X-rays are generated. Where the collimator edges are straight, the resulting beam that subtends the detector will have curve to it and hence software will have to be provided to compensate for the curvature. However, where the collimator includes curved edges that are selected appropriately, the beam subtending the detector will not include curvature and less software processing will be required.